The present invention relates to the field of electrical leads suitable for being implanted within living tissues, and more particularly to implantable electrical leads having relatively low electrical resistance which may be used in conjunction with cardiac pulse generators, neural stimulators, implantable sensors, and the like.
An implantable cardiac pulse generator, referred to generically as a cardiac pacer, or pacemaker, is a small, sealed electronic pulse generator that is used to treat irregular heart rhythms. In general, such pacers provide minute electrical stimuli to a heart when needed to speed up unnaturally slow heart rates.
An implantable cardiac defibrillator is a moderately sized, electronic pulse generator that is used to treat patients that are at risk from suffering lethal arrhythmias, most notably ventricular fibrillation. Ventricular fibrillation is a heart rhythm that typically results in death within several minutes. The defibrillator is used to provide large electrical stimuli when needed to interrupt the lethal arrhythmia and re-establish a life sustaining heart rhythm. Such pulse generators are typically packaged in sealed containers that are usually implanted subcutaneously in the thorax or abdomen of the heart patient. These devices monitor cardiac activity and deliver electrical pulses of appropriate intensity whenever needed. The energy supplied by a pulse generator is conducted along an electrically conductive cardiac lead from the pulse generator directly to the heart.
The pulse generator is commonly powered by a battery located inside the sealed container which is not intended to be replaceable. The amount of electrical energy stored in the battery generally determines the operational life of the pulse generator. Although the battery is very efficient at storing electrical energy, the battery life, and hence the operational life of the pulse generator, is usually less than ten years. The battery depletion is in part due to energy delivered to the heart, to energy consumed by the resistance of the electronic circuitry of the pulse generator and cardiac leads, and to self-discharge of the battery over time.
Each time an electrical pulse is delivered to the heart, some of the energy output of the battery is consumed by the cardiac leads as I.sup.2 R heat, where "I" represents the current through the cardiac lead, and "R" represents the electrical resistance of the cardiac lead. The I.sup.2 R losses represent wasted energy which provides no useful purpose. In an attempt to maximize the service life of the pulse generator, the lead materials and the geometries of the lead materials are chosen to minimize the electrical resistance of the cardiac leads while providing a lead that can withstand the rigors of exposure to repetitive stress.
The lead which electrically connects the pulse generator to the heart may be attached to the inner surface of the heart, the endocardium, or to the outer surface of the heart, the epicardium. Regardless of where the lead is attached to the heart, the lead is mechanically flexed with every heart beat. Every flexure of the lead creates stress within it. Since a typical heart rate is 60 beats per minute, the heart beats millions of times in a single year, and the lead is stressed with each beat.
Unlike a pulse generator, which is replaced when the battery is depleted, the lead is not normally replaced. A youthful patient who receives an implant may hopefully use the same lead or leads for decades. For this reason the materials comprising the lead should have excellent mechanical fatigue resistance.
Materials known as having low electrical resistance, such as copper or silver, are not well suited for use as a conductor in a cardiac lead because they cause tissue reactions and readily form oxides which may ultimately result in the fracture of a lead constructed with such material. A further disadvantage of copper and silver is that they have very poor resistance to repeated stress. Core copper wire, multi-stranded copper wire and even tinsel copper wire would poorly withstand the repeated, reversing stresses the heart would impose on cardiac leads comprised of copper wire. Leads of such constructions tend to fail after only a relatively small number of flexions, much as a paper clip breaks after being bent a few times.
Spring materials made of non-oxidizing, corrosion resistant alloys having good fatigue resistance perform much better mechanically as cardiac lead conductors than do conductors comprised of copper, silver, or their alloys. That is why for decades lead manufacturers have been using spring materials for the conductors in their leads. Examples of suitable conducting materials include stainless steel, such as Elgin Wire Co., Elgiloy, MP-35N, and titanium and titanium alloys.
Generally, in the construction of cardiac leads, the conducting wire core is coiled to form a tight helix composed of many individual coils, similar to an extension spring. The helical construction greatly lowers the mechanical stress to which the material comprising the wire core would otherwise be subjected by the beating heart. Though this construction provides long lasting leads, the electrical resistance of leads manufactured of spring steel or titanium is relatively high primarily due to the resistivity of the material comprising the wire core.
Low resistance leads are important both for pacing and defibrillation. DBS (drawn brazed strands) and DFT (drawn filled tubing) provide a cardiac lead having both reasonable fatigue resistance and electrical resistance. DBS and DFT are examples of two structures which combine low resistance, poor mechanical materials with the high resistance, spring materials of conventional leads. An example of DBS wire includes six strands of wire made of MP-35N that are brazed by a central silver core. An example of drawn filled tubing may include tubing fabricated of MP-35N and which is filled with silver. The silver, copper or other electrical conductor significantly reduces the lead resistance, but the silver or copper included in these leads also present several drawbacks. Chief among them are the toxicity and low fatigue life of copper and silver. There are several materials, as previously mentioned that are well suited to be pacing or defibrillation lead conductors. However, these materials unfortunately have high resistivities of about 100 micro-ohm-cm, as opposed to resistivities of about 1 micro-ohm-cm for copper or silver.
The electrical resistance of a typical cardiac lead is about 100 ohms. Cardiac leads with larger diameter conductors, high helix pitches (a measure of the number of coils per unit of axial length of the lead), multiple conductors, or smaller helical internal diameters, may have electrical resistances of about 10 ohms. However, such leads also tend to have reduced flex life.
A goal in the field of cardiac lead technology is to provide a cardiac lead having an electrical resistance of less than 1 ohm. Low resistance cardiac leads would provide a pulse generator with an increased service life because such leads would reduce I.sup.2 R losses. Thus, more energy stored in the battery which powers the pulse generator would be able to be delivered to the heart because less energy would be wasted as heat. Therefore, it may be appreciated that there is a need for a cardiac lead having low electrical resistance, as well as good resistance to the repeated stresses to which a cardiac lead is exposed.